Dosimetric scintillating screen detector for charged particle radiotherapy quality assurance (qa)

ABSTRACT

An apparatus and method are provided for performing Quality Assurance of complex beams of penetrating radiation inside a patient. A detector with a transverse scintillating screen images the radiation inside a tissue phantom with high spatial resolution. The scintillator is comprised of a mixture of two or more scintillators emitting different spectra of light and having different characteristic responses as a function of the beam LET value. The optics relaying the scintillation output have variable transmission with wavelength, further shaping the spectrum of light transmitted to the imaging sensor which also has spectrally varying sensitivity. Parameters of the scintillator construction, the optics, and the imaging sensor are chosen so the output of the composite detector is proportional to a characteristic of the input beam, for example the dose deposited as a function of depth inside the tissue phantom.

REFERENCE TO PENDING PRIOR PATENT APPLICATION

This patent application claims benefit of pending prior U.S. ProvisionalPatent Application Ser. No. 61/602,301, filed Feb. 23, 2012 by Steven M.Ebstein for DOSIMETRIC SCINTILLATING SCREEN DETECTOR FOR CHARGEDPARTICLE RADIOTHERAPY QA (Attorney's Docket No. EBSTEIN-6 PROV), whichpatent application is hereby incorporated herein by reference.

FEDERALLY SPONSORED RESEARCH

This invention was made with U.S. Government support under Grant5R44CA103610-03 awarded by the National Cancer Institute. The U.S.Government has certain rights in the invention.

FIELD OF THE INVENTION

The present invention pertains to a device and a method for monitoringthe spatial characteristics of a beam of penetrating radiation and, moreparticularly, to quality assurance in particle therapy, namely, ensuringthat the dose delivered matches the intended treatment plan.

BACKGROUND OF THE INVENTION

The goal of radiotherapy is to irradiate unhealthy, usually cancerous,tissue or tumors with the aim of killing that tissue while sparing, tothe extent possible, the surrounding healthy tissue. This requires thatthe patient be irradiated with one or more radiation beams that arecarefully positioned and shaped in order to precisely deliver theintended quantity of radiation to a specific target volume. Indimensions transverse to the nominal direction of the beam, this shapingis often performed with metal apertures or multi-leaf collimators thatdefine the profile of the beam. The beams are often directed fromdifferent angles at the patient, thereby increasing the radiation dosein the intersecting regions. A treatment plan is generated for eachirradiation session of a patient that describes the various beam angles,shapes, energies, and quantity of radiation delivered along with theassociated maps showing the radiation dose intended for the patient.

With high-energy X-ray beams, the technique of spatially shaping theradiation dose with multi-leaf collimators and multiple beams directedfrom multiple angles is known as Intensity Modulated Radiation Therapyor IMRT.

Charged particles, typically protons but also heavier ions, are alsoused for delivering spatially complex radiotherapy treatments. Theparticle accelerators typically produce a narrow, or pencil, beam thatcan be focused, collimated, or deflected with magnetic optics, althoughsometimes the beam is broadened in size or angle by scattering thepencil beam. With particle beams, the shaping can also be effected alongthe beam direction by varying the energy distribution of the chargedparticles, as the penetration distance and dose distribution along theparticles' tracks depends on their energy. For a monoenergetic ion beamin a homogeneous medium, the absorbed dose versus depth describes acurve known as a pristine Bragg peak as seen at 12 in FIG. 1.

From entrance into the body until close to the penetration depth, thedelivered dose, which is proportional to the average linear energytransfer (LET) given in keV/μm due to ionization by the beam, risesslowly along the plateau of the curve. Near the final depth, the LETincreases rapidly to a peak, then falls to zero rapidly at the end ofthe track of the ensemble of particles in the beam, thereby increasingthe radiation dose in a narrow volume. The LET value is typically acharacteristic of a beam as a function of depth. It is related to theenergy loss per unit length dE/dx per unit density, typically given inMeV cm²/g, of each charged particle. dE/dx is a function of the particleenergy.

In this specification, the term beam generally refers to a penetratingradiation beam. However, it may either refer to a radiation beamproduced by an accelerator, a narrow pencil beam at the output of thebeam delivery system, a set of pencil beams delivered within a shorttime, i.e., a composite beam, or a broad beam restricted by a metalaperture or collimator. A treatment field generally refers to acomposite or broad beam with a set of particle energies and a definedspatial distribution delivered within a short time. If either theenergies or spatial distribution or both are changed, that generallyrefers to a new treatment field.

By irradiating with beams having different energies, or by modulatingthe beam energy as it is applied, the dose distribution can be shaped todeposit maximal dose in an extended volume containing the tumor whiledelivering reduced dose to the surrounding healthy tissue. An example,shown at 14 in FIG. 1, uses a beam with a set of different energies andfluences arranged to produce a depth-dose curve with a flat maximum,denoted as a spread-out Bragg peak or SOBP. The energy of particles inthe beam can be reduced by inserting absorbing elements, typicallyplastic, upstream of the patient. The energy reduction depends on thethickness of the absorber, so a set or range of energies can be producedwith a spatially varying thickness to the absorber or by rotating awheel with multiple thickness steps and temporally modulating the beamenergy.

FIG. 1 highlights one of the differences between radiotherapy usingx-rays or charged particles. For an equivalent dose at the depth in themiddle of the SOBP, the x-ray beam deposits more dose in front of thatdepth and after the tail of the Bragg peak, represented by the areabetween the x-ray dose curve 10 and the charged particle SOBP dose curve14. This ability to limit the dose to healthy tissue outside of a tumoris a major motivation for delivering radiotherapy with chargedparticles.

Charged particle beams can also be scanned across the patient if thebeam is small enough and scanning magnets are provided to steer the beamin angles away from the central axis of the system. This technique isgenerally termed pencil beam scanning or PBS. PBS treatments can bedelivered without metal apertures or collimators if the scanning magnetscan sufficiently define the transverse shape of the composite beam.Typically, more than one radiation field will be applied from a singledirection with different energies or energy spectra used for thedifferent fields, thereby shaping the delivered dose at differentdepths.

Several different forms of PBS are used to deliver radiotherapy. Atechnique known as Intensity Modulated Proton Therapy or IMPT allows thebeam intensity to vary as a function of lateral position by modulatingthe beam current or modulating the scanning rate. In some scenarios, thebeam position is continually scanned and in others it is translated to aseries of fixed positions where the beam is turned on for a period oftime, a technique called spot scanning. For this approach, the magneticscanning of the beam defines the shape of the irradiated volumetransverse to the beam direction. An approach, which is termed eitherwobbling or uniform scanning, has the beam scan a given area, typicallya regular region like a circle or rectangle, and deliver a uniform doseover that area. A fixed aperture or multi-leaf collimator is positionedbetween the beam source and the patient to restrict the transversedistribution of radiation.

In all of these various forms of radiotherapy, the standard forpositioning and shaping the beam relative to the patient is typicallyless than one millimeter across a field of regard that can range up to40 cm in diameter. This requires that the radiation pattern be deliveredwith precise shaping and positioning of the beam(s) over those scales,as well as in precise amounts so the dose absorbed by the patientmatches the prescription of the treatment plan. This requires carefulcalibration and measurement of the delivery system and the treatmentplan, i.e., the beam(s) which will be delivered to the patient, aprocess referred to as quality assurance or QA. QA measurements areroutinely performed that reassure the practitioners that the correctradiation dose will be delivered in the correct amount over the correctspatial distribution to the patient. These measurements are performedwith a variety of different radiation detectors.

With the complex series of radiation fields that are used inradiotherapy, especially in charged particle therapy, it is critical tomake measurements which predict the absorbed dose inside of a patient,not merely in a plane just outside of the patient. One way this isaccomplished is by placing absorbing material, a tissue phantom, inbetween the beam source and the detector so the radiation incident onthe detector corresponds to the radiation incident on the correspondinglocation(s) inside of the patient. The absorbing material is typicallywater (since the body is mostly made up of water), sheets of plasticwith water-equivalent absorbance, or more complex phantoms with multiplematerials.

There exist several different detectors used for performing QA of theseradiotherapy beams. Film has been used since the beginnings ofradiotherapy to view the beam as delivered to the patient. Various othertechnologies have been used, including scintillating screen based portalimagers, flat panel devices consisting of a sheet of scintillatingcrystals atop an array of semiconductor detectors, ionization chambersas single devices, stacked along the beam direction, or arrayedtransverse to the beam direction, multi-wire proportional chambers,gaseous electron multiplier (GEM) detectors, and other detectors, etc.Typically, these detectors are one part of the overall system andprocedure for performing QA.

Requirements for these detectors include stability, robustness,linearity, accuracy, and precision, i.e., signal to noise ratio (SNR),spatial resolution, and dosimetric accuracy. The detector output shouldbe translatable into a measure of the effective dose delivered to thepatient. In addition, marketability factors such as the cost andusability of the detector output are important for the usefulness of adetector for QA applications.

An ideal QA detector would measure the delivered dose inside a tissuephantom with sufficient spatial resolution to validate the treatmentplan. In practice, current detectors fall short in one or more respects.Some of the detectors employed, like film and detectors using thinscintillating screens or sheets of crystal, have spatial resolution thatexceeds the requirement for beam positioning and shaping. However, theirdetection signal is not linear with the fluence or LET of the radiationbeam, which is required in order to predict the corresponding dose tothe patient.

The light output of some scintillators will saturate when a clinicallyuseful fluence is incident. Inorganic scintillators or phosphors such asP11 (ZnS:Ag), P20 (Zn, CdS:Ag), P43 (Gd₂O₂S:Tb), P46 (Y₃Al₅O₁₂:Ce), andP47 (Y2SiO₅:Ce) have responses that are typically linear with fluence.However, scintillator light output is generally not linear with the LETof the incident particles. A plot of the light output versus depth hasthe same general shape as the Bragg peak. However, the ratio of the peakvalue to the value at a point on the plateau differs from the ratio ofLET values calculated using accurate physical models or that is measuredby a standard detector such as an ion chamber. When the ratio of lightoutput is less than the ratio of LET or absorbed dose, this is termedquenching. The physical explanation is that an abundance of excitedatoms provide alternate pathways for excited species to relax by meansother than photon emission.

Scintillating screens are used, in conjunction with radiation therapy,for absorbed dose measurement (as described, for example, by J. M.Schippers, S. N. Boon and P. van Luijk, “Applications in Radiationtherapy of a scintillating screen viewed by a CCD camera,” Nucl Instr.and Meth. Λ 477, pp. 480-85 (2002), and S. N. Boon, thesis, “Dosimetryand quality control of scanning proton beams” (1998), available onlineat http://www.ubsug.nl/eldoc/dis/science/s.n.boon/, and referencestherein, all of which are incorporated herein by reference). They arealso useful for monitoring beam delivery in real time (S. M. Ebstein,High Resolution Proton Beam Monitor, U.S. Pat. No. 7,515,681). However,as those references show, knowledge of the particle beam energy spectrumand careful calibration of the detector response is required to makeaccurate dosimetric measurements.

Other detectors such as ion chambers are quite linear in theirmeasurement of the fluence and LET, so they are quite accurate inpredicting absorbed dose. However, due to their size and complexity,they cannot be made into arrays that can measure the incident radiationdistribution with the required resolution.

Due to the limitations of these detectors, a complex system is requiredto perform and evaluate the QA measurements. Typically, a series ofrepresentative measurements are made with detectors that have highspatial resolution but poor dosimetric accuracy or detectors withlimited spatial resolution but good dosimetric accuracy. The response ofthe detectors to the delivered radiation beams is calculated from amodel of the beam delivery and detection system and compared to theactual measurements. This approach is described in C. Brusasco and B.Marchand, Device And Method For Particle Therapy Monitoring AndVerification, U.S. Patent Application Pub. No. US 201110248188 A1.

However, this approach cannot give complete confidence that thedelivered beam matches the plan due to reduced dimensionality of themeasurements. In addition, the work required to construct the models andperform the calculations of the expected measurements adds to the costand complexity of performing QA.

There exists a need for improved detectors, especially for chargedparticle therapy QA. The beam delivery systems are complex instrumentsand are required to produce a series of beams with complex shapes anddifferent energy spectra. Current detectors and systems cannot directlymeasure the delivered dose as a function of depth with sufficientaccuracy or spatial resolution to quickly and easily perform the QAtask. A detector which combines the high spatial resolution of eitherfilm or a scintillating screen detector and the dosimetric accuracy ofan ion chamber would be very useful in performing QA of complexradiotherapy treatment plans, especially for IMPT.

SUMMARY OF THE INVENTION

In accordance with preferred embodiments of the present invention, anapparatus and method are provided for performing QA of complex beams ofpenetrating radiation that produce a spatially varying dose distributioninside of a patient. The apparatus provides for inserting tissue phantommaterial in between the beam source and a detector. The detectorincorporates a scintillator disposed transverse to the nominal directionof the beam and one or more high resolution imaging sensors in opticalcommunication with the scintillator for generating an image of the beamdistribution across the scintillator. The scintillator is comprised of amixture of two or more scintillators emitting different spectra of lightand having different characteristic responses as a function of the beamLET value. The means of optical communication of the scintillationoutput, i.e., the optics, have variable transmission with wavelength,further shaping the spectrum of light transmitted to the imaging sensor.The imaging sensor has sensitivity that varies across the spectrum oflight input. Parameters of the scintillator construction, the optics,and the number and types of imaging sensors are chosen so the output ofthe composite detector is proportional to a characteristic of the inputbeam. In one embodiment, the characteristic is the dose deposited as afunction of depth inside the tissue phantom. In another embodiment, thescintillator is disposed in a thin, planar screen that is orientedperpendicular to the central beam axis. In another embodiment, thescintillator output is communicated to multiple sensors whose outputsare combined numerically to produce an output proportional to thedesired characteristic.

In one preferred form of the present invention, there is provided anapparatus for determining the spatial distribution and intensity ofpenetrating radiation beams characterized generally by a propagationdirection inside a body, the apparatus comprising:

a tissue phantom disposed between the radiation source and a radiationdetector;

a scintillating screen disposed behind the tissue phantom for emittinglight in response to the radiation comprising a mixture of at least twoscintillators wherein each scintillator has a different characteristicresponse and a different spectral output;

a means of optical communication of the scintillator output to at leastone imaging sensor wherein the means of optical communication has anonuniform spectral transmission; and

at least one imaging sensor in optical communication with thescintillating screen for providing a high resolution imaging sensoroutput indicative of the spatial distribution and intensity of theradiation beam wherein the imaging sensor has a nonuniform spectralsensitivity;

wherein the composition of the scintillator, the means of opticalcommunication, and imaging sensors are selected so as to comprise asystem wherein the imaging sensor output is proportional to acharacteristic of the radiation beam incident on the scintillator ateach measurement position accessible with the tissue phantom.

In another preferred form of the present invention, there is provided amethod for determining the spatial distribution and intensity ofpenetrating radiation beams characterized generally by a propagationdirection inside a body, the method comprising:

providing a tissue phantom between the radiation source and a radiationdetector;

providing a scintillating screen behind the tissue phantom for emittinglight in response to the radiation comprising a mixture of at least twoscintillators wherein each scintillator has a different characteristicresponse and a different spectral output;

providing a means of optical communication of the scintillator output toat least one imaging sensor wherein the means of optical communicationhas a nonuniform spectral transmission; and

providing at least one imaging sensor in optical communication with thescintillating screen for providing a high resolution imaging sensoroutput indicative of the spatial distribution and intensity of theradiation beam wherein the imaging sensor has a nonuniform spectralsensitivity;

wherein the composition of the scintillator, the means of opticalcommunication, and imaging sensor are selected so as to comprise asystem wherein the imaging sensor output is proportional to acharacteristic of the radiation beam incident on the scintillator ateach measurement position accessible with the tissue phantom.

In another preferred form of the present invention, there is provided amethod for determining the spatial distribution and intensity ofpenetrating radiation beams characterized generally by a propagationdirection inside a body, the method comprising:

providing a tissue phantom between the radiation source and a radiationdetector;

providing a scintillating screen behind the tissue phantom for emittinglight in response to the radiation comprising a mixture of at least twoscintillators wherein each scintillator has a different characteristicresponse and a different spectral output;

providing a means of optical communication of the scintillator output toan imaging sensor wherein the means of optical communication has anonuniform spectral transmission; and

providing an imaging sensor in optical communication with thescintillating screen for providing a high resolution imaging sensoroutput indicative of the spatial distribution and intensity of theradiation beam wherein the imaging sensor has a nonuniform spectralsensitivity;

wherein a first composition of the scintillator, a first means ofoptical communication, and a first imaging sensor are selected so as tocomprise a first system wherein the imaging sensor output isproportional to a characteristic of the radiation beam incident on thescintillator at each measurement position accessible with the tissuephantom; and

performing characteristic measurements with a penetrating radiation beamand the first system;

wherein one or more of (a) the composition of the scintillator, (b) themeans of optical communication, or (c) the imaging sensor is adjusted soas to comprise a second system wherein the imaging sensor output isproportional to a characteristic of the radiation beam incident on thescintillator at each measurement position accessible with the tissuephantom with greater accuracy than with the first system.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing features of the invention will be more readily understoodby reference to the following detailed description, taken with referenceto the accompanying drawings, in which:

FIG. 1 depicts the relative radiation dose from a photon beam, apristine proton beam (Bragg peak), and a spread out Bragg peak (SOBP)using protons composed of multiple pristine proton Bragg peaks ofdifferent energies. The area between the curves is the excess dose froma photon beam.

FIG. 2 shows a schematic depiction of a system for delivering aspatially complex radiation beam.

FIG. 3 shows a schematic depiction of the QA apparatus including tissuephantom material, a scintillating screen, and optics relaying the imageto an imaging sensor.

FIG. 4 shows the luminous efficiency of a scintillator versus dE/dx.

FIG. 5 shows the spectral content of scintillator output for variousscintillators.

FIG. 6 shows exemplary spectral transmission of compound lenses.

FIG. 7 shows exemplary spectral transmission of colored glass filters.

FIG. 8 shows exemplary spectral response of an imaging sensor.

FIG. 9 shows depth dose measurements with a Markus ion chamber andseveral scintillators.

FIG. 10 shows exemplary Relative Biological Effectiveness (RBE) versusLinear Energy Transfer (LET).

FIG. 11 shows an apparatus which records multiple images of thescintillator.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

In accordance with preferred embodiments of the present invention, thespatial distribution of a beam of penetrating radiation is imaged with ascintillator. The term “penetrating radiation,” as used herein, and inany appended claims, refers both to particles with mass, such asprotons, as well as to photons, i.e., to electromagnetic radiation suchas x-rays or gamma rays. Moreover, in the case of massive particles, theparticles are typically charged, such as protons or heavier atomic ions,however, neutrons or other electrically neutral particles may also bedetected, and their beams imaged within the scope of the presentinvention.

Referring to FIG. 2, a system for delivering a spatially complexradiation beam is designated generally by 20. The system consists of aradiation source, 21, and a means of setting the energy of theradiation, 22, setting the fluence, and means for shaping the transversespatial distribution of the radiation, 23. In the case of chargedparticle beams, the shaping means can comprise a scatterer whichscatters a unidirectional beam into a beam with a wide angular spread incombination with a fixed or variable collimator or aperture whichdefines the shape of the beam. Alternatively, the shaping means cancomprise two orthogonal scanning magnets which can steer the angle of apencil beam, 24, from a virtual point located near the magnets across afield of regard, 25.

Referring to FIG. 3, a QA apparatus consists of a tissue phantom 31disposed between the beam shaping means and the detector 40 which imagesthe radiation beam. There is provision for varying the effective depthwithin the phantom, e.g., by varying the thickness of the phantommaterial, where the detector is located. For example, the tissue phantommay consist of a series of sheets of a polymer material, e.g., awater-equivalent plastic. Sheets may be added or removed so as to varythe effective depth within the phantom where the detector images thebeam. In order to simulate measurement at depth within a body, thedetector and tissue phantom are typically translated away from the beamsource so as to keep the position of the initial surface of the phantomat the same distance from the beam source.

Alternatively, the tissue phantom may comprise a water tank. With thebeam directed downwards, as is possible with many radiotherapy deliverysystems which comprise a rotating gantry, the water level in the tankcan be continually varied. As water is added to the tank, the water tankand detector can be translated downward, away from the beam source, soas to preserve the distance from the beam source to the top of the waterlayer, thereby simulating measurement at depth within a body. If thewater tank comprises a bellows-like apparatus, the thickness of thewater layer can be varied with a beam that is not vertical.

The detector comprises a thin scintillating screen 41, denoted thescintillator, at least one imaging sensor 44, and a means of opticalcommunication of the scintillation light from the screen to the sensor,42 and 43. In order for the detector output to match a characteristicsuch as the absorbed dose, the scintillator comprises a mixture of twoor more scintillating materials comprising a thin layer bound to asupporting substrate, typically a thin plastic film such as Mylar. Sincethe output of a single material will not typically match the desiredcharacteristic as a function of depth within the tissue phantom, thescintillator composition comprises a mixture of materials whose combinedoutput, in combination with the rest of the apparatus, can match thedesired characteristic.

For example, it is known that many scintillators such as GSO, LYSO,Y3Al5O12:Ce (P46), Gd2O2S:Tb (P43), and Y2SiO5:Ce (P47) exhibitquenching with increasing dE/dx or LET. That is, their light output perdeposited energy, dL/dE, decreases monotonically with increasing dE/dxsuch as occurs within the Bragg peak. However, there are scintillatorssuch as NaI(Tl), CsI(Tl), ZnS:Ag (P11) and (Zn, Cd)S:Ag (P20) which showanti-quenching behavior, i.e., dL/dE is not monotonic and increases withdE/dx for some values. Examples of both quenching and anti-quenchingbehavior are shown in R. B. Murray and A. Meyer, Scintillation Responseof Activated Inorganic Crystals to Various Charged Particles, Phys. Rev.Vol 122, No. 3, pp. 815-826 (1961), Y. Koba, H. Iwamoto, K. Kiyohara, T.Nagasaki, G. Wakabayashi, Y. Uozumi, N. Matsufuji, “ScintillationEfficiency of Inorganic Scintillators for Intermediate-Energy ChargedParticles,” Progress in Nuclear Science And Technology, Vol. 1,p.218-221 (2011), S. Safai, S. Lin, and E. Pedroni, “Development of aninorganic scintillating mixture for proton beam verification dosimetry,”Phys Med Biol. 2004 Oct. 7;49(19):4637-55 and S. Safai, InorganicScintillating Mixture And A Sensor Assembly For Charged ParticleDosimetry, U.S. Pat. No. 8,080,801. FIG. 4 shows the luminous efficiencyof several scintillators as a function of dE/dx.

One can generate an output signal that is proportional to a desiredcharacteristic by combining the outputs from different scintillators inways detailed below. It is apparent that for these common scintillators,different characteristic responses such as quenching and anti-quenchingbehavior are accompanied by different spectral outputs. Therefore, thedetection and combination of the signals from each scintillator mustaccount for their varying spectral outputs. In order to do so, severalparameters must be chosen so that they work in combination to produce asignal that is linear in the desired characteristic. These parametersinclude (a) the scintillator composition; (b) the spectral transmissionof the optics; and (c) the number of and spectral sensitivity of imagingsensors which detect the scintillated light.

In a preferred embodiment of the invention, the scintillating screencomprises a mixture of a quenching and an anti-quenching inorganicscintillator comprising a substrate comprised of a thin sheet ofsupporting material, e.g., polyester sheet, and a thin layer of thescintillating material ranging from 10-1000 mg/cm² in thickness. Inaddition to the scintillating material, a polymer material is typicallyincluded as a binding agent to dimensionally stabilize the scintillatormixture and adhere the scintillating layer to the substrate thatsupports it. The substrate, the scintillator mixture and any bindingagents will affect the characteristics of the scintillation lightoutput. The substrate may be clear or partially reflective. Theinorganic scintillators are typically opaque and will scatter or absorbsome or all of the scintillation light output from the portions of thescintillating layer proximal to the source of the radiation beam. Avariety of binding agents are known to the art, including a variety ofsingle component and dual component optical cements and epoxies whichcure by heat, UV light, or chemically. These binding agents are selectedbecause of one or more characteristics including their cost, ease ofuse, their ability to form a uniform scintillating layer, their opticalclarity, and their ability to withstand radiation without darkening,yellowing, or otherwise suffering radiation damage.

In a preferred embodiment of the invention, the light from thescintillator is communicated to the imaging sensor by a mirror and acompound lens close to the imaging sensor. The mirror, designated thefold mirror, directs light from the scintillator to the side so theimaging sensor can be located outside of the radiation beam. The lensforms a high resolution image of the scintillating screen on the imagingsensor so there is an accurate, high resolution sampling of theradiation beam as it is incident on the scintillator. Typically, boththe fold mirror and the lens will have a nonuniform spectral effect onthe scintillation light transmitted to the sensor. The reflectance ofthe fold mirror, especially when used at a 45 degree angle, varies withwavelength over the range of visible scintillators, and lenstransmission varies with wavelength, as well.

In a preferred embodiment of the invention, the imaging sensor comprisesa thermoelectrically cooled CCD detector.

Safai et al. have disclosed a mixture scintillator which has onecomponent with a quenching characteristic and one with an anti-quenchingcharacteristic, which in combination produces a signal that has aquenching characteristic different than either component alone. QA ofthe treatment would be simpler if the signal is proportional to thedeposited dose. However, the spectral output of each scintillatorcomponent is typically not identical. Thus, using the typical measure ofscintillator efficiency, output power per input power (W/W), is notsufficient to predict the detection signal. The signal will depend onthe interaction of the spectral content of the scintillator output withthe spectral transmission of the optics and the spectral sensitivity ofdetectors which receive the light from the mixture scintillator. Asignal proportional to deposited dose may result with certain optics anddetectors but not with others.

The spectral content and quantity of the scintillation light is afunction of several multiplicative factors within the linear range ofthe scintillator and the sensor. First, each scintillating component hasa characteristic spectrum S_(n)(λ)as well as an areal density ρ_(n).Most important, its luminous efficiency dL/dE=E_(n)(LET, F), i.e., theunit light output is a function of LET and the overall fluence F. Theoutput is attenuated by a factor Λ_(n)(λ)due to the self-absorption andscattering by the scintillation layer which depends on the scintillatordensity and the characteristics of the binder such as the binder arealdensity and transmission. The imaging sensor will typically have asensitivity Q(λ) that reflects the effective quantum efficiency of thesensor, as a function of wavelength, in addition to a gain factor thatdepends on the electronic amplification and digitization of thephotodetection signal. Finally, the spectral transmission T(λ) of theoptics transmitting the scintillation light to the sensor alsodetermines the detected signal.

Thus, the detection signal D(LET, F) per deposited energy equals

${D\left( {{LET},F} \right)} = {\underset{n}{\Sigma}{\int{{\; {\lambda\rho}_{n}}{S_{n}(\lambda)}{E_{n}\left( {{LET},F} \right)}{A_{n}(\lambda)}{Q(\lambda)}{T(\lambda)}}}}$

For this equation, we have assumed that the characteristics of thescintillating screen are uniform across its area. While that is onlytrue with a given accuracy—scintillating screen thickness can typicallyvary by several percent across a screen tens of cm in size—the variationis generally independent of LET and multiplies D(LET) uniformly as afunction of position or angle which can be calibrated by a flat fieldcorrection. In addition, evidence from Boon is that the spectral contentof the scintillation output does not vary with LET for some commonscintillators.

In order for the detection signal to match a characteristic of theradiation beam, the densities of the components of the mixture must beengineered so that the detection signal matches that characteristic, fora given spectral transmission T(λ) and spectral sensitivity Q(λ). For adosimetric detector, the desired characteristic is that the detectionoutput is proportional to the energy deposited, i.e., D(LET, F) isessentially independent of the LET value and the fluence. For a twocomponent scintillator, that can be accomplished if one component showsdecreasing output with LET (quenching), the other shows increasingoutput with LET (anti-quenching), and the proportions of the twocomponents are such that the combination has uniform output with LET.Within a range of densities, the proportions should reduce to a ratio ofthe density of the two components.

However, it may be difficult to design that ratio in advance. Factorssuch as A_(n)(λ) factor are not well understood and hard to control, asit can be difficult in practice to precisely control the absolute valueof ρ_(n) deposited on the substrate and its uniformity. Factors such asQ(λ) are typically average values for sensors tested with a specificinput beam of light. A particular sensor with different conditions ofthe input beam may yield a measurably different efficiency as a functionof wavelength due to interference effects resulting from the thicknessof a semiconductor sensor, for example.

It is possible to tailor the specifications of the optics to compensatefor such variations. FIG. 5 shows the considerable variation in thespectral content of the light output of different inorganicscintillators. Thus, the relative contribution of different componentsof a scintillator mixture to the detection signal can be adjusted byvarying the spectral throughput of the optics. The optics transmissionfunction T(λ) can be modified by including an element in the optics withvariable transmission or reflection as a function of wavelength. In thecase of the fold mirror, different reflective coatings such as aluminum,protected aluminum, and various dielectric coatings have non-uniformspectral reflectivity as a function of wavelength. In the case of thelens, different optical glasses have non-uniform spectral transmission.In addition, coatings applied to the glass can have non-uniform spectraltransmission. FIG. 6 shows exemplary transmission curves for compoundlenses. Furthermore, a color filter can be used with a lens to enhanceor diminish various portions of the spectrum.

Colored filters such as Wratten filters, Schott colored glass, and Hoyacolored glass can be used for this purpose. FIG. 7 shows exemplarytransmission curves for colored glass filters. In addition, numerousfilters used for color temperature adjustment, color correction, andlight balancing are available. These latter filters are typicallydesigned to correct the effective color (black body) temperature fromartificial (tungsten or fluorescent) light to that of sunlight, or viceversa. They often come in a series which can apply smaller or largercolor corrections depending on the input lighting or desired output.

It is also possible to employ variable optical components in order tomodify the spectral response and hence the degree to which the detectionsignal matches the desired characteristic. For example, variable opticalfilters are available with changeable passbands or transmissionproperties that are modified by changing the temperature of the filter,the voltage across a liquid crystal transmissive element, or by someother means.

Another consideration is that properties of the system components mayvary over time due to gradual modification of their properties by aging,slow chemical reactions, or radiation damage. For example, radiationexposure can color or darken optical glasses and optical cements such asthe polymer binder in the scintillator, or it can introduce crystaldefects into the scintillating material that change the efficiency ofthe scintillator. The performance of optical coatings can change overtime if they absorb moisture. Absorbing dyes in colored filters canchange, chemically, over time to alter their transmission properties.Any of these changes will affect the elements of the right-hand side ofthe equation describing the detection signal given above.

Some of these variations in output over time do not affect the basicability of the system to measure the appropriate characteristic. Aspectrally uniform reduction of the overall signal, e.g., due to acolorless (gray) darkening of an optical element, effectively multipliesthe detection signal equation by a constant factor independent of thewavelength λ. This multiplicative or scale factor can be calibrated bymeasuring the detection response to a radiation beam whose fluence andenergy spectrum are known from measurements with other detectors.

Other variations can affect the ability of the system to measure theappropriate characteristic. Variations that change the relativecontributions of the various components in the scintillation mixture tothe detection signal may change whether the system linearly measures theappropriate characteristic of the radiation. Those variations maycomprise a differential change in (a) the output of the scintillatorcomponents or (b) a change in how that scintillator output istransmitted or detected. For example, one of the optical elements couldundergo a colored darkening of a material due to radiation damage whichwould reduce the contribution from a first scintillator component morethan from a second scintillator component.

A variation in one element of the equation can be compensated for bychanging one or more of the other elements of the system. For example,the colored darkening from the previous example could be compensated forby adding a filter which reduces transmission from the second componentmore than from the first component. Alternatively, if radiation damagereduces output from a first component of the scintillator more rapidlythan from a second component of the scintillator, a filter could beadded that preferentially transmits the spectral output of the firstcomponent relative to the second component so as to compensate.

It may be difficult to predict the elements of the right-hand side ofthe detection equation with sufficient accuracy. As noted, thescintillator output can be quite sensitive to the scintillator mixture,including the binder, and the layer thickness. Moreover, the values ofsome elements of the equation may have significant uncertainties due toprocess variation. There may also be practical considerations wheretechnical specifications for some components are unavailable and it maybe impractical to measure those specifications of the components. Inthose cases, some assumptions about the specifications of those elementscan be made in order to generate an initial design of the system. Thedesign would include selection of the mixture proportions and thevarious components.

The initial design can be tested by performing calibration measurementswith a well-characterized radiation beam, preferably monoenergetic, andobserving the detection signal as a function of depth inside the tissuephantom. Due to the uncertainties, the signal may not match the desiredcharacteristic to a sufficient accuracy. In that case, the design can bemodified by changing either the scintillator composition, the opticalsystem, or the imaging sensor. For example, if dose is the desiredcharacteristic and it is found that the detection signal using atwo-component scintillator exhibits a quenching characteristic, anotherscintillator could be fabricated with the fraction of the quenchingphosphor reduced relative to the initial design.

Another approach involves adding or changing an optical filter thatenhances the contribution of one component of the scintillator.Referring to the previous example, a filter could be added that hashigher transmission for the spectral content of the anti-quenchingcomponent than for the spectral content of the quenching component. Afilter with the necessary transmission ratio for the two components canadjust the output of the system so it matches the desired characteristicresponse. In the case of dosimetry, the adjusted system would notexhibit quenching or anti-quenching.

The images from the imaging sensor must undergo some processing in orderto render the image value at each pixel proportional to the lightincident on the sensor. It is known to the art that two point correctionconsisting of subtraction of a bias value and multiplication of theresult by a gain factor is generally required to linearize the output ofan imaging detector. Moreover, these two point corrections typicallyvary across the imaging sensor and from one copy of a sensor to anotherso these correction factors are typically specific to a particulardetector and take on different values for each pixel of the sensor. Wethus speak of the bias and gain images or arrays that are used toprocess the raw sensor output, since they vary for each pixel.

A series of calibrations are employed to measure the two pointcorrection factors, i.e., recording the sensor response when noradiation is incident (a dark exposure) or the response to a uniformradiation pattern (flat field). In some cases, the sensor output isnon-linear with the underlying characteristic, e.g., the intensity orfluence of the incident radiation. In these cases, a non-lineartransformation of the sensor signal is performed after its response hasbeen carefully calibrated. In what follows, the resulting signalsdescribed are those after the imaging sensor output has been processedto produce a signal proportional to the incident light.

In accordance with a preferred embodiment of the invention, thescintillator comprises a mixture of two components, by weight, of 25%P11 and 75% P43 in a layer 120 mg/cm² on a thin polyester sheet with apolymer binder. The resulting screen has transverse dimensions severalcentimeters or greater in order to simultaneously measure an area withclinical significance, i.e., an area that surrounds a tumor. The opticscomprise a 45 degree first surface minor of enhanced aluminum coated onglass and a Canon EF camera lens with 50 mm focal length. The imagingsensor comprises a Kodak KAF two-phase CCD employing a transparent gate.FIG. 8 shows the typical sensitivity of this CCD as a function ofwavelength. With this composition, the image output is proportional tothe radiation dose as measured by a Markus ion chamber.

Note that this composition of the scintillator comprises ananti-quenching blue phosphor, P11, and a quenching yellow-greenphosphor, P43. This choice of a scintillator mixture with significantlydifferent color outputs enables the detection properties to be easilytuned with spectrally-varying components that enhance blue versus yellowand green or vice versa.

A favorable aspect of this composition is that it does not use toxicmaterials. Recently, the desire to reduce the use of hazardous materialssuch as lead in solder has led to regulatory requirements for electronicand other instruments which do not use such materials. This trend isexemplified by designation of ROHS (reduction of hazardous substances)qualified components, which are now required in some jurisdictions.Other scintillators such as P20 or (Zn, Cd)S:Ag include cadmium, ahazardous material.

FIG. 9 shows the results of measurements with a Markus ion chamber, aquenching scintillator (P46), and two mixture scintillators irradiatedby a narrow energy spectrum proton beam with varying thicknesses of aplastic absorber. The “Mix1” scintillator was designed with twoscintillator components whose proportions and thickness, together withthe optics and imaging sensor, produce an output that matches the doseas measured by the Markus chamber. The Mix1 scintillator is comprised,by weight, of 25% P11 and 75% P43. The “Mix2” scintillator was designedwith a different proportion of the same two scintillator components tobe somewhat quenching. The Mix2 scintillator is comprised, by weight, of12.5% P11 and 87.5% P43.

A scaling factor was applied to the P46, Mix1 and Mix2 measurements sothe plateau region (starting at 0 cm depth) of all 4 data sets lie onthe same curve. As the curves approach the Bragg peak, they separate dueto whether they exhibit quenching or not, relative to the ion chamberresponse. As is evident from FIG. 9, the Mix1 output matches the Markuschamber output very well and the Mix2 output exhibits some quenching,though not as much as the P46 scintillator.

Measurements such as those shown in FIG. 9 indicate how well thedetection measurements match a characteristic, in this case defined bythe ion chamber response. Although the detection measurements with theMix2 scintillator do not match the characteristic, the method describedabove could be used to adjust the detection response, namely, a filtercould be added with higher transmission for blue light than yellow-greenlight. That would have the effect of enhancing the anti-quenchingcomponent of the signal.

In addition to dosimetry, other characteristics of the depositedradiation may be of interest. For example, the efficacy of radiationtherapy is a function of the relative biological efficiency (RBE) of aparticular radiation beam. Different types of radiation—photons orprotons or heavier ions—have different RBE values and the RBE may be afunction of the LET value. FIG. 7 shows the RBE values as a function ofLET for various probabilities of cell survival. So, it may be desirableto measure a detection signal that is appropriately weighted to matchthe RBE of the radiation beam for a given probability of cell survival.If the scintillator components have a different light output per unitenergy depending on LET, then scintillator mixtures can be produced withlight outputs that match the appropriate RBE of the beam for a givenprobability of cell survival.

A properly designed system has the detection signal proportional to thedesired characteristic. Especially as treatment modes incorporate verynarrow pencil beams, the incident fluence may increase to the pointwhere the luminous efficiency (light output) E_(n)(LET, F) of one ormore scintillator components is not linear with the fluence F butsaturates with increasing fluence. If there is significant differentialnon-linearity of the scintillator components, the detection signal mayonly be linear with the characteristic to the required accuracy for alimited range of fluence. Since pencil radiation beams typically have aGaussian intensity profile, the fluence varies as a function ofposition, and the linear measurement may only be achievable over aportion of the beam spot.

The system design can be tuned so the linear range corresponds with theclinically relevant range of fluence. In this fashion, measurement of acharacteristic of the beam that is limited to a range of fluence can belinearized with the multiple component scintillator. Furthermore, achangeable optical filter, either by replacement or use of a variablefilter, can be used to dynamically tune the optical system. Thus onefilter (or setting of a variable filter) can be used for one range offluence and another filter (or setting of a variable filter) can be usedfor a different range of fluence.

The discussion above shows how a custom spectral filter can compensatefor the detailed characteristics of a particular mixture scintillator.Moreover, changing the filter can shift the linear range of the systemwith respect to fluence. Changing the filter can also compensate foraging effects of the scintillator.

In a preferred embodiment of the system, multiple, spectrally distinctimages of the scintillator are formed and combined to produce an outputproportional to the desired characteristic. The contributions of thevarious scintillator components can be separated by spectrally filteringthe scintillation light and forming multiple images of the beamscintillation. Then, a detection signal can be produced by an arbitrarylinear combination of the spectrally separated images. Methods are knownto the art of forming multiple images with substantially non-overlappingspectral content. Examples include video cameras with a color separationprism and 3 CCDs to separately record red, green, and blue (RGB)components.

In another example, shown in FIG. 11, a dichroic mirror 45 separates theblue output of P11 and the yellow-green output of P43 and dual imagesare recorded by using two separate imaging sensors 44 and 47.Alternatively, the dual images could be spatially multiplexed onto asingle imaging sensor. The two images can be spatially registered andco-added with a weighting factor that produces the best fit to linearizethe output with respect to the desired characteristic, e.g., thewater-equivalent dose. Before co-adding the images, standard two-pointcorrection (spatially variable bias and gain) or a more complexcorrection is applied to each image so it is a linear representation ofthe incident light.

More complicated detection signals can be synthesized from the two (ormore) spectral images than simply co-adding the images with a singleweighting factor. If one or more of the scintillator components exhibitssaturation with increasing fluence, a non-linear function of thespectral images can be constructed that linearizes the output withrespect to fluence, as well as the desired characteristic. This can beimplemented by calibrating the response and modeling the output signalat each pixel as a polynomial function of the corresponding pixel ineach spectrally distinct image. Alternatively, a lookup table that isindexed by the corresponding value from each separate image can be used.

Modifications

The described embodiments of the invention are intended to be merelyexemplary and numerous variations and modifications will be apparent tothose skilled in the art. All such variations and modifications areintended to be within the scope of the present invention as defined inthe appended claims.

What is claimed is:
 1. An apparatus for determining the spatialdistribution and intensity of penetrating radiation beams characterizedgenerally by a propagation direction inside a body, the apparatuscomprising: a tissue phantom disposed between the radiation source and aradiation detector; a scintillating screen disposed behind the tissuephantom for emitting light in response to the radiation comprising amixture of at least two scintillators wherein each scintillator has adifferent characteristic response and a different spectral output; ameans of optical communication of the scintillator output to at leastone imaging sensor wherein the means of optical communication has anonuniform spectral transmission; and at least one imaging sensor inoptical communication with the scintillating screen for providing a highresolution imaging sensor output indicative of the spatial distributionand intensity of the radiation beam wherein the imaging sensor has anonuniform spectral sensitivity; wherein the composition of thescintillator, the means of optical communication, and imaging sensorsare selected so as to comprise a system wherein the imaging sensoroutput is proportional to a characteristic of the radiation beamincident on the scintillator at each measurement position accessiblewith the tissue phantom.
 2. The apparatus as set forth in claim 1,wherein the characteristic of the radiation beam is the dose.
 3. Theapparatus as set forth in claim 1, wherein the characteristic of theradiation beam is the response of an ion chamber.
 4. The apparatus asset forth in claim 1, wherein the characteristic of the radiation beamis the relative biological efficiency.
 5. The apparatus as set forth inclaim 1, wherein the scintillator comprises a mixture of at least twoinorganic scintillators.
 6. The apparatus as set forth in claim 1,wherein the scintillator components are chosen from a group of inorganicscintillators including P11, P20, P43, P46, and P47.
 7. The apparatus asset forth in claim 1, wherein the scintillator components include P11and P43.
 8. The apparatus as set forth in claim 1, wherein thescintillator comprises a planar screen disposed perpendicular to thepropagation direction of the radiation.
 9. The apparatus as set forth inclaim 1, wherein the means of optical communication comprise a foldminor and a compound lens.
 10. The apparatus as set forth in claim 1,wherein the imaging sensor comprises a CCD.
 11. The apparatus as setforth in claim 1, wherein the tissue phantom comprises polymer sheets.12. The apparatus as set forth in claim 1 wherein the tissue phantomcomprises water.
 13. The apparatus as set forth in claim 1 wherein themeans of optical communication include at least one filter withnonuniform spectral transmission.
 14. The apparatus as set forth inclaim 1 wherein the means of optical communication include at least onefilter with variable nonuniform spectral transmission.
 15. The apparatusas set forth in claim 1 wherein the means of optical communicationinclude a dichroic mirror which separates the scintillated light intotwo substantially non-overlapping spectral bands and forms two spatiallydistinct images on the at least one imaging sensor.
 16. The apparatus asset forth in claim 1 wherein the characteristic of the radiation beamincludes the fluence.
 17. A method for determining the spatialdistribution and intensity of penetrating radiation beams characterizedgenerally by a propagation direction inside a body, the methodcomprising: providing a tissue phantom between the radiation source anda radiation detector; providing a scintillating screen behind the tissuephantom for emitting light in response to the radiation comprising amixture of at least two scintillators wherein each scintillator has adifferent characteristic response and a different spectral output;providing a means of optical communication of the scintillator output toat least one imaging sensor wherein the means of optical communicationhas a nonuniform spectral transmission; and providing at least oneimaging sensor in optical communication with the scintillating screenfor providing a high resolution imaging sensor output indicative of thespatial distribution and intensity of the radiation beam wherein theimaging sensor has a nonuniform spectral sensitivity; wherein thecomposition of the scintillator, the means of optical communication, andimaging sensor are selected so as to comprise a system wherein theimaging sensor output is proportional to a characteristic of theradiation beam incident on the scintillator at each measurement positionaccessible with the tissue phantom.
 18. The method as set forth in claim17, wherein the characteristic of the radiation beam is the dose. 19.The method as set forth in claim 17, wherein the characteristic of theradiation beam is the response of an ion chamber.
 20. The method as setforth in claim 17, wherein the characteristic of the radiation beam isthe relative biological efficiency.
 21. The method as set forth in claim17, wherein the scintillator comprises a mixture of at least twoinorganic scintillators.
 22. The method as set forth in claim 17,wherein the scintillator components are chosen from a group of inorganicscintillators including P11, P20, P43, P46, and P47.
 23. The method asset forth in claim 17, wherein the scintillator components include P11and P43.
 24. The method as set forth in claim 17, wherein thescintillator comprises a planar screen disposed perpendicular to thepropagation direction of the radiation.
 25. The method as set forth inclaim 17, wherein the means of optical communication comprise a foldminor and a compound lens.
 26. The method as set forth in claim 17,wherein the imaging sensor comprises a CCD.
 27. The method as set forthin claim 17, wherein the tissue phantom comprises polymer sheets. 28.The method as set forth in claim 17 wherein the tissue phantom compriseswater.
 29. The method as set forth in claim 17 wherein the means ofoptical communication include at least one filter with nonuniformspectral transmission.
 30. The method as set forth in claim 17 whereinthe means of optical communication include at least one filter withvariable nonuniform spectral transmission.
 31. The method as set forthin claim 17 wherein the means of optical communication include adichroic mirror which separates the scintillated light into twosubstantially non-overlapping spectral bands and forms two spatiallydistinct images on the at least one imaging sensor.
 32. The method asset forth in claim 17 wherein the characteristic of the radiation beamincludes the fluence.
 33. A method for determining the spatialdistribution and intensity of penetrating radiation beams characterizedgenerally by a propagation direction inside a body, the methodcomprising: providing a tissue phantom between the radiation source anda radiation detector; providing a scintillating screen behind the tissuephantom for emitting light in response to the radiation comprising amixture of at least two scintillators wherein each scintillator has adifferent characteristic response and a different spectral output;providing a means of optical communication of the scintillator output toan imaging sensor wherein the means of optical communication has anonuniform spectral transmission; and providing an imaging sensor inoptical communication with the scintillating screen for providing a highresolution imaging sensor output indicative of the spatial distributionand intensity of the radiation beam wherein the imaging sensor has anonuniform spectral sensitivity; wherein a first composition of thescintillator, a first means of optical communication, and a firstimaging sensor are selected so as to comprise a first system wherein theimaging sensor output is proportional to a characteristic of theradiation beam incident on the scintillator at each measurement positionaccessible with the tissue phantom; and performing characteristicmeasurements with a penetrating radiation beam and the first system;wherein one or more of (a) the composition of the scintillator, (b) themeans of optical communication, or (c) the imaging sensor is adjusted soas to comprise a second system wherein the imaging sensor output isproportional to a characteristic of the radiation beam incident on thescintillator at each measurement position accessible with the tissuephantom with greater accuracy than with the first system.
 34. The methodas set forth in claim 33 wherein the adjustment comprises one of (a)addition of an optical filter or (b) replacement of an optical filter.35. The method as set forth in claim 33 wherein the characteristic ofthe radiation beam includes the fluence.